Alongside magnetic resonance tomography (MR), in recent years, positron emission tomography (PET) has also become increasingly widespread in medical diagnosis. Whereas MR is an imaging method for displaying structures and slice images in the interior of the body, PET enables the visualization and quantification of metabolic activities in-vivo.
PET uses the particular properties of positron emitters and positron annihilation in order to determine the function of organs or cell areas quantitatively. Thereby, before the examination, the patient is administered appropriate radiopharmaceuticals that are marked with radionuclides. As they decay, the radionuclides emit positrons that interact with an electron after a short distance resulting in so-called annihilation. Two gamma quanta are then created and fly apart from each other in opposite directions (offset by 180°). The gamma quanta are detected by two opposing PET detector modules inside a specific timeframe (coincidence measurement) from which the location of annihilation is determined at a position on the connecting line between these two detector modules.
For detection purposes, the PET detector module must generally cover a major part of the length of the gantry arc. It is divided into detector elements with side lengths of a few millimeters. On detecting a gamma quantum, each detector element generates an event record that specifies the time and the detection location, that is the corresponding detector element. This information is transferred to a fast logic unit and compared. If two events coincide within a maximum time interval, it is assumed a gamma decay process has occurred on the connecting line between the two associated detector elements. The reconstruction of the PET image is performed with the aid of a tomography algorithm, the so-called back projection.
Since MR systems operate with high magnetic fields, it is necessary to use materials compatible therewith within these systems. In particular, when designing PET detectors in combined PET-MR systems, it is necessary to ensure that the detectors are insensitive to magnetic fields.
With combined PET-MR systems, it is known to use lutetium oxyorthosilicate (LSO) as a scintillation crystal for converting the gamma quanta into light and avalanche photodiodes (APDs) for detecting the light. The APDs are connected to preamplifiers. A ring of PET detectors of this kind is arranged inside an MR appliance. This enables MR and PET data records to be recorded simultaneously.
In particular with the commonly used semiconductor amplifiers and semiconductor detectors, the gain depends upon the temperature. Since the components are subjected to temperature fluctuations during operation, cooling is necessary. The temperature of the amplifiers and photodiodes can be controlled by supplying cooled air. When using air with a constant temperature, the temperature of the amplifiers results from the balance of the generated heat and the heat emitted through air via the surfaces of the amplifiers. The cooling can be used in the same fashion for other parts of the detection system.
However, APDs are not subjected to only temperature fluctuations due to their operation. In particular, the proximity to the gradient coil and the excitation coil in the MR system resulting from the compact design represents a heat source acting on the APDs from the outside. The temperature of a gradient coil is typically between 20 and 80° C. during operation. These temperature differences also affect the APDs and hence their gain. The effects of this heat source can only be controlled with difficulty by way of air cooling.
The sensitive electronics in the electronic circuits associated with PET detectors also have to be protected from overheating.